1. Field of the Invention
This invention lies in the field of polymer chemistry in which the polymers formed are suitable for coating biosensors. The coatings act to decrease the impedance at the sensor""s electrode and thereby enhance the signal during in vivo placement of the sensor.
2. Description of Related Art
Biosensors are small devices that use biological recognition properties for selective analysis of various analytes or biomolecules. Typically, the sensor will produce a signal that is quantitatively related to the concentration of the analyte. To achieve a quantitative signal, a recognition molecule or combination of molecules is often immobilized at a suitable transducer which converts the biological recognition event into a quantitative response.
A variety of biosensors have been developed for use with numerous analytes. Electroenzymatic biosensors use enzymes to convert a concentration of analyte to an electrical signal. Immunological biosensors rely on molecular recognition of an analyte by, for example, antibodies. Chemoreceptor biosensors use chemoreceptor arrays such as those of the olfactory system or nerve fibers from the antennules of the blue crab Callinectes sapidus to detect the presence of amino acids in concentrations as low as 10xe2x88x929 M. For a review of some of the operating principles of biosensors, see Bergveld, et al., ADVANCES IN BIOSENSORS, Supplement 1, p. 31-91, Turner ed., and Collison, et al., Anal. Chem 62:425-437 (1990).
Regardless of the type of biosensor, each must possess certain properties to function in vivo and provide an adequate signal. First, the elements of the biosensor must be compatible with the tissue to which it is attached and be adequately shielded from adjacent tissues such that allergic or toxic effects are not exerted. Further, the sensor should be shielded from the environment to control drift in the generated signal. Finally, the sensor should accurately measure the analyte in the presence of proteins, electrolytes and medications which may interfere.
The prototype biosensor is the amperometric glucose sensor. There are several reasons for the wide ranging interest in glucose sensors. In the healthcare arena, glucose sensors are useful for glucose monitoring of patients with diabetes mellitus. Additionally, a working glucose sensor is required for the development of a closed loop artificial pancreas with an implanted insulin pump. A commercial interest focuses on sensors that can be used to monitor fermentation reactions in the biotechnology arena. From a scientific standpoint, interest is driven by the availability of a very robust enzyme, glucose oxidase, which can be used to monitor glucose, as well as the desire to develop model sensors for a wide variety of analytes.
Any amperometric glucose sensor or any oxido-reductase enzyme that uses O2 as a co-substrate and is designed for subcutaneous or intravenous use requires both an outer membrane and an anti-interference membrane. The requirement of two distinct membranes is due to the fundamental nature of the sensor as well as the environment in which the measurement is made.
A glucose sensor works according to the following chemical reaction (Equation 1): 
In this reaction, glucose reacts with oxygen in the presence of glucose oxidase (GOX) to form gluconolactone and hydrogen peroxide. The gluconolactone further reacts with water to hydrolyze the lactone ring and produce gluconic acid. The H2O2 reacts electrochemically as shown below (Equation 2):
H2O2xe2x86x92O2+2exe2x88x922H+xe2x80x83xe2x80x83(II) 
The current measured by the sensor/potentiostat (+0.5 to +0.7 v oxidation at Pt black electrode) is due to the two electrons generated by the oxidation of the H2O2. Alternatively, one can measure the decrease in the oxygen by amperometric measurement (xe2x88x920.5 to xe2x88x921 V reduction at a Pt black electrode).
The stoichiometry of Equation 1 clearly demonstrates some of the problems with an implantable glucose sensor. If there is excess oxygen for Equation 1, then the H2O2 is stoichiometrically related to the amount of glucose that reacts at the enzyme. In this case, the ultimate current is also proportional to the amount of glucose that reacts with the enzyme. If there is insufficient oxygen for all of the glucose to react with the enzyme, then the current will be proportional to the oxygen concentration, not the glucose concentration. For the sensor to be a true glucose sensor, glucose must be the limiting reagent, i.e. the O2 concentration must be in excess for all potential glucose concentrations. For a number of conditions, this requirement is not easily achieved. For example, the glucose concentration in the body of a diabetic patient can vary from 2 to 30 mM (millimoles per liter or 36 to 540 mg/dl), whereas the typical oxygen concentration in the tissue is 0.02 to 0.2 mM (see, Fisher, et al., Biomed. Biochem. Acta. 48:965-971 (1989). This ratio in the body means that the sensor would be running in the Michaelis Menten limited regime and would be very insensitive to small changes in the glucose concentration. This problem has been called the xe2x80x9coxygen deficit problemxe2x80x9d. Accordingly, a method or system must be devised to either increase the O2 in the GOX membrane, decrease the glucose concentration, or devise a sensor that does not use O2.
Several approaches to solving the deficit problem have been attempted in the past. The simplest approach is to make a membrane that is fully O2 permeable, with no glucose permeability and mechanically perforate it with a small hole that allows glucose to pass. Here the differential permeability is defined by the ratio of the small hole area to the total membrane area. Two significant problems with this method are first that reproducibly making small holes is difficult and second and more serious, the O2 permeability is a strong function of the thickness of the membrane and thickness is difficult to control in mass production. Microporous membranes (U.S. Pat. No. 4,759,828 to Young et al., incorporated herein by reference) have also been tried with limited success. Another problem with both the perforated membrane approach and the microporous membrane approach is that the sensor electrodes and the enzyme layer are exposed to body fluids. Body fluids contain proteins that coat the electrodes leading to decreased sensitivity of the sensor and enzymes (proteases) that can digest or degrade the sensor active enzyme.
Another approach to the oxygen deficit problem is described by Gough (U.S. Pat. No. 4,484,987, incorporated herein by reference). The approach uses a combination membrane with discrete domains of a hydrophilic material embedded in a hydrophobic membrane. In this case, the membrane is not homogenous and manufacturing reproducibility is difficult. Physical properties of the membrane are also compromised. In a similar manner, Gough (U.S. Pat. No. 4,890,620, incorporated herein by reference) describes a xe2x80x9ctwo dimensionalxe2x80x9d system where glucose diffusion is limited to one dimension while the oxygen diffusion is from both dimensions. This sensor is extremely complicated and manufacturing on a large scale is expected to be difficult.
Several other groups have used a homogenous membrane of a relatively hydrophobic polyurethane and reported good results. See, for example, Shaw, et al., Biosensors and Bioelectronics, 6:401-406 (1991); Bindra, et al., Anal. Chem 63:1692 (1991); and Schichiri, et al., Horm. Metab. Resl. Suppl. Ser., 20:17 (1988). In classical diffusion experiments with these membranes, however, the glucose diffusion is extremely small. It is believed that the ability of these polyurethane layers to allow glucose diffusion is due to micro cracks or micro holes in these materials when applied as membranes.
Still others have developed homogeneous membranes with both hydrophilic and hydrophobic regions to circumvent the oxygen deficit problem. See, Allen et al., U.S. Pat. Nos. 5,284,140 and 5,322,063, the disclosures of each being incorporated herein by reference. These patents describe acrylic and polyurethane systems, respectively. Both of the membranes have hydrophilic and hydrophobic moieties in the molecule leading to limited control of oxygen and glucose permeabilities.
The key to stable, high sensitivity enzyme biosensors is that the sensor output must be limited only by the analyte of interest, not by any co-substrates or kinetically controlled parameters such as diffusion. In order to maximize the output current (Equation 2) of the biosensor, oxygen diffusion should be as large as possible while maintaining oxygen excess at the reaction surface. Since the normal concentration of O2 in the subcutaneous tissue is quite low, maximization of the O2 diffusion coefficient is desirable.
The membrane systems described in the literature as cited above attempt only to circumvent the oxygen deficit problem by reducing the amount of glucose diffusion to the working electrode of the biosensor. There is a need for the membrane to have physical stability and strength, adhesion to the substrate, processibility (ability to be synthesized/manufactured in reasonable quantities and at reasonable prices), biocompatibility, ability to be cut by laser ablation (or some other large scale processing method), and compatibility with the enzyme as deposited on the sensor.
Another one of the problems with implantable biosensors occurs as a result of xe2x80x9croad blockxe2x80x9d type interference. This problem is encountered when the outermost layer of the biosensor has some hydrophobic characteristics. These characteristics result in the accumulation of plasma proteins on the surface of the electrode after only short periods of direct contact with body fluids. The hydrophobic regions of the sensor surface are believed to denature the proteins resulting in large deposits of protein mass. The deposits then affect the sensor""s performance through a physical interference in a xe2x80x9croad blockxe2x80x9d type of effect. The protein deposition is a gradual process which creates a non-uniform, non-predictable diffusion path for the analyte to the sensor. Moreover, the effect on the sensor is a cascading type in which the protein deposits dissipate the normal voltages applied to the electrodes (i.e., the deposits increase the capacitance of the system). The resultant requirement for higher voltages to offset the increased capacitance increases the noise, ultimately compromising the validity of the sensor""s output.
Other problems are also associated with implantable sensors having hydrophobic regions at the sensor""s surface. In particular, subcutaneous tissue contains substantial amounts of lipid vesicles. By implanting a biosensor directly into tissue, a portion of the sensor may be implanted directly into, or flush against a very hydrophobic lipid region. This also limits the aqueous environment which is required around the sensor""s electrodes.
What is needed in the art are new coatings for implantable sensors which are extremely hydrophilic and provide a substantial and uniform aqueous flow around the sensors. Quite surprisingly, the present invention provides such coatings and sensors equipped with those coatings.
The present invention provides methods for reducing the electrode impedance of implantable biosensors by coating the surface of the biosensor with a uniform hydrogel which allows unimpeded water movement around the sensor. The surface coatings are compositions which are biocompatible and are capable of water uptake of at least 120% of their weight, more preferably at least 200% of their weight. Upon the uptake of water, the hydrogels used in the present invention will also swell and provide a layer of water around the electrodes to which the hydrogels are attached.
In one group of embodiments, the hydrogels can be prepared from (a) a diisocyanate, (b) a hydrophilic polymer which is a hydrophilic diol, a hydrophilic diamine, or a combination thereof, and optionally, (c) a chain extender.
The present invention also provides silicon containing compositions which are biocompatible and suitable for coating a biosensor. The compositions are polymers which are formed into membranes and can be prepared from: (a) a diisocyanate, (b) a hydrophilic polymer which is a hydrophilic diol, a hydrophilic diamine, or a combination thereof, (c) a siloxane polymer having functional groups at the chain termini, and optionally, (d) a chain extender.
The membranes prepared from the above components will have a glucose diffusion coefficient of from about 1xc3x9710xe2x88x929 cm2/sec to about 200xc3x9710xe2x88x929 cm2/sec, a water pickup of at least about 25% and a ratio of Doxygen/Dglucose of about 5 to about 200.
In certain preferred embodiments, the functional groups present in the siloxane polymer are amino, hydroxyl or carboxylic acid, more preferably amino or hydroxyl groups. In other preferred embodiments, the hydrophilic polymer is a poly(ethylene)glycol which is PEG 200, PEG 400 or PEG 600. In still other preferred embodiments the diisocyanate is a isophorone diisocyanate, 1,6-hexamethylene diisocyanate or 4,4xe2x80x2-methylenebis(cyclohexyl isocyanate) and the chain extender is an alkylene diol, an alkylene diamine, an aminoalkanol or a combinations thereof.
In particularly preferred embodiments, the diisocyanate is 1,6-hexamethylene diisocyanate, the hydrophilic polymer is PEG 400 or PEG 600 and is present in an amount of about 17 to about 32 mol % (relative to all reactants), and the siloxane polymer is aminopropyl polysiloxane having a molecular weight of about 2000 to about 4000 and is present in an amount of about 17 to about 32 mol % (relative to all reactants).
The present invention further provides an implantable biosensor for measuring the reaction of an analyte, preferably glucose, and oxygen, the biosensor having a biocompatible membrane as described above. The present invention further provides implantable biosensors for measuring a variety of analytes, the biosensor having a coating as described above.